Implantable cardioverter-defibrillators (ICDs) are used to provide various types of therapy to a cardiac patient, including, for example cardioversion and/or defibrillation. These devices consist of a hermetic housing implanted into a patient and connected to at least one defibrillation electrode and with at least one other electrode e.g., a patch-type electrode, a housing- or can-based electrode, a surface-type electrode, and a stent-based electrode) thereby defining a therapy vector between various pairs of said electrodes. The housing of the ICD contains electronic circuitry for monitoring the condition of the patient's heart, usually through sensing electrodes, and also contains the battery, high voltage circuitry and control circuitry to generate, control and deliver the defibrillation shocks. Typically, one or more specialized defibrillation-type or other transvenous leads are connected to circuitry within the ICD and extend from the housing to one or more defibrillator electrodes proximate the heart. The housing of the ICD may include one or more defibrillation electrodes configured on the exterior of the housing. One example of an ICD is disclosed in U.S. Pat. No. 5,405,363 to Kroll et al., the disclosure of which is hereby incorporated by reference.
One important parameter for the effective operation of an ICD device is the defibrillation electrode impedance. This impedance is indicative of the positioning and integrity of the defibrillation leads and/or electrodes. Electrode impedance is also related to the defibrillation threshold for a given patient used in setting the energy levels for defibrillation shocks for that patient. Successful cardiac defibrillation depends on the amount of energy applied to the cardiac tissue by the electrical defibrillation shock, and the energy of the defibrillation shock is dependent on the electrode impedance of the defibrillation electrodes through which the defibrillation shock is delivered.
Determining the impedance between defibrillation electrodes is used in different ways when implanting and operating an ICD. One use is to allow a physician to verify that the defibrillation leads and/or electrodes have not shifted after an initial placement. Another use is to permit the physician to adjust waveform durations in the event of a significant impedance change. Still another use is to confirm the viability and settings appropriate for a defibrillation shock prior to delivering the defibrillation shock. Thus, it can be seen that accurate knowledge of the electrode impedance is important both during implantation and operation of an ICD device.
Presently, ICD devices periodically measure the impedance across the defibrillation leads by using a low voltage monophasic or alternating square wave pulse on the order of 10 volts. Most ICD devices use a low voltage monophasic pulse that is generated from the battery, rather than the high voltage capacitors that are used to generate and deliver a defibrillation shock. This is done both to keep the test shock at a level that is below the pain or perception level that may be felt by a patient, and to minimize the drain on the battery in order to periodically supply these test shocks.
With a normal defibrillation shock, the current passed through the defibrillation electrodes is on the order of ten amperes and several hundred volts and many charge carriers in the cardiac tissue are recruited to carry this current. When the cardiac tissue is subjected to a lower current pulse, fewer charge carriers are recruited to carry the lower current. As a result, the impedance of the cardiac tissue in response to a lower current pulse increases significantly. For example, a forty ohm (Ω) defibrillation pathway might have an apparent impedance of over 120Ω with a lower voltage and correspondingly lower current pulse. This differential behavior of cardiac tissue in response to different amounts of current is discussed by Brewer J E, Tvedt M A, Adams T P, and Kroll M W in Low Voltage Shocks Have a Significantly Higher Tilt of the Internal Electric Field Than Do High Voltage Shocks, PACING AND CLINICAL ELECTROPHYSIOLOGY Vol. 18, p. 214 (January 1995), the disclosure of which is hereby incorporated by reference.
Because this differential behavior of cardiac tissue is known, current ICD devices using a low voltage pulse to measure the impedance of defibrillation electrode will generate a measured value that can be as much as three times greater than the actual defibrillation impedance encountered for a high voltage, high current defibrillation shock. Consequently, current ICD devices typically divide the impedance measured in response to a low voltage test pulse by some kind of “fudge” factor (e.g., 2 or 3) to estimate the actual impedance. Unfortunately, the fudge factor is not consistent with all types of leads, electrodes, patients, or changing electrolyte concentrations. Thus, significant errors are often introduced that may yield inconsistent impedance measurements.
One approach to reducing the errors induced by the use of low voltage test shocks for measuring defibrillation electrode impedance is described in U.S. Pat. No. 6,104,954 to Blundsden. In one embodiment, a square wave generator is described to generate a test pulse train of approximately 50 V and 100 Khz. While this approach would somewhat improve the accuracy of the impedance measurement, unfortunately this embodiment is completed impractical for an ICD device as the continuous power requirement to implement this kind of square wave test pulse would be 50 W, an amount which is well above any continuous power supply that can be provided by current defibrillation battery technology of an ICD device.
In another embodiment, Blunsden describes the use of a test pulse train having a higher voltage shock in the range of defibrillation voltages that is delivered from the high voltage capacitors in the ICD as a shorter test shock pulse train for purposes of measuring defibrillation electrode impedance. The approach has the advantage of testing not only the defibrillation electrodes, but also the operation of the high voltage switches used to generate biphasic pulses that are typically used for defibrillation shocks. While this approach has the added advantage of exercising the high voltage switches and can address some of the errors induced by the use of low voltage test shocks, the approach introduces the possibility of unwanted shocks in the event of a failure of the high voltage switches or heightened sensitivity of the patient to larger voltage shocks. The approach also requires an increase in the drain on the battery required in order to periodically charge the high voltage capacitors to deliver these defibrillation range shocks for measuring the defibrillation electrode impedance where the vast majority of the energy required to charge the high voltage capacitors is ultimately wasted.
Defibrillation strength shocks (approximately ten amperes and several hundred volts) are extremely painful and cannot be given to conscious patients. Accurately measuring the defibrillation electrode impedance is important to effective operation of an ICD device. There is an unfulfilled need to accurately measure the actual impedance between defibrillation electrodes while minimizing or eliminating the sensation of pain felt by the patient and not adversely affecting the overall performance of the ICD device.